Functional Optical coherence tomography (OCT) extensions provide additional contrast that partially mitigate the missing specificity of OCT. Doppler OCT and OCT angiography are among the functional modalities that are closest to their clinical translation. The latter provides microvascular contrast without the need of dye administration, thus allowing for screening and frequent treatment monitoring. The principle is based on observing signal decorrelation or motion contrast due to moving red blood cells within vessels. There exist a multitude of methods to extract this decorrelation signature. The easiest method to be used with swept source OCT is based on speckle decorrelation [16]. All methods have in common, that dense sampling is needed to achieve good vascular contrast. The highest sensitivity is achieved when signal decorrelation between successive tomograms is assessed. Taking several tomograms at the same location however comes at the price of measurement time, which should be kept short for retinal imaging. Otherwise motion artifacts cause degradation of contrast and resolution.
Phase-resolved Doppler optical coherence tomography (PR-D-OCT) is a non-invasive imaging technique that provides depth-resolved quantitative knowledge about motion with high resolution and at high speed [1]. Applied to functional tissue imaging, the blood velocity of selected vessels can be determined together with their cross sectional diameter, which permits calculation of the volumetric blood flow. Knowledge of tissue perfusion provides valuable information about tissue health or treatment progression. PR-D-OCT is however only sensitive to the projection of the velocity along the illumination beam, and therefore requires knowledge of the Doppler angle, a, between vessel orientation and optical axis, to obtain the absolute velocity. At this point it should be mentioned that apart from PR-D-OCT also other methods exist to determine the axial flow velocity component or in general the velocity component in direction of illumination and detection such as resonant Doppler OCT [26], or joint time frequency domain OCT [25]. All methods that measure the axial velocity component share the same disadvantage of missing knowledge about the Doppler angle.
Several absolute velocity methods have been developed in an attempt to address this issue. First approaches employed the tomographic information readily available with OCT to determine α [2-4]. While being a valid approach for rather steep vessels, an accurate evaluation of the velocity becomes challenging in vascular plexuses close to perpendicular to the optical axis because of the precision required on α. Recently, a method demonstrated a direct calculation of the blood flow from en face cross sections [5]. It advantageously uses the fact that the dependence of the velocity and vessel cross section in the en face plane on cos(α) cancels each other for the flow calculation. Again, an accurate value is only obtained for steep vessels like the one present in the optic nerve head (ONH) of the human eye [6-8]. Other methods rely on the determination of the 3D velocity vector by measuring the same sample point under different angles with several beams and corresponding detection channels with detailed knowledge of the probing beam geometry. Dual-beam bidirectional OCT [9] employs two different directions and allows the determination of the absolute velocity by knowledge of β, the angle between illumination plane and vessel orientation in the en face plane. This technique is particularly suitable for posterior pole blood flow assessment. Indeed, multi-beam methods rely on the exact superposition of beams, a condition that becomes critical to fulfill with increasing number of beams and the limited optical quality of the eye. The flow is later calculated by multiplying the absolute velocity with the vessel cross section, usually obtained from a separate fundus camera [10]. Acquisitions with large β require high precision knowledge of its value. Moreover, the method ultimately fails for β reaching 90°. We previously demonstrated an extension of the method's flexibility by rotating the illumination plane with a Dove prism to an angle parallel to the vessel orientation [11]. Under such condition, the sensitivity of the velocity calculation on β is low. Still, for more complex vascular structures this condition cannot be met for every vessel at the same time. Also, because of patient motion, the scanning trajectory can deviate from the ideal and expected position. To account for β would therefore require a live fundus camera or a registration algorithm [12].